The background description provided herein is for the purpose of generally presenting the context of the disclosure. Work of the presently named inventors, to the extent the work is described in this background section, as well as aspects of the description that may not otherwise qualify as prior art at the time of filing, are neither expressly nor impliedly admitted as prior art against the present disclosure.
The use of positron emission tomography is growing in the field of medical imaging. In positron emission tomography (PET) imaging, a radiopharmaceutical agent is introduced into the object to be imaged via injection, inhalation, or ingestion. After administration of the radiopharmaceutical, the physical and bio-molecular properties of the agent cause it to concentrate at specific locations in the human body. The actual spatial distribution of the agent, the intensity of the region of accumulation of the agent, and the kinetics of the process from administration to its eventual elimination are all factors that may have clinical significance. During this process, a positron emitter attached to the radiopharmaceutical agent will emit positrons according to the physical properties of the isotope, such as half-life, branching ratio, etc.
The radionuclide emits positrons, and when an emitted positron collides with an electron, an annihilation event occurs, wherein the positron and electron are combined. Most of the time, an annihilation event produces two gamma rays (at 511 keV) traveling at substantially 180 degrees apart.
In order to be able to reconstruct the spatio-temporal distribution of the radio-isotope via tomographic reconstruction principles, each detected event will need to be characterized for its energy (i.e., amount of light generated), its location, and its timing. By detecting the two gamma rays, and drawing a line between their locations, i.e., the line-of-response (LOR), one can determine the likely location of the original disintegration. While this process will only identify a line of possible interaction, by accumulating a large number of those lines, and through a tomographic reconstruction process, the original distribution can be estimated. In addition to the location of the two scintillation events, if accurate timing (within a few hundred picoseconds) is available, a time-of-flight (TOF) calculation can add more information regarding the likely position of the event along the line. Limitations in the timing resolution of the scanner will determine the accuracy of the positioning along this line. The collection of a large number of events creates the necessary information for an image of an object to be estimated through tomographic reconstruction.
PET imaging systems use detectors positioned across from one another to detect the gamma rays emitting from the object. Typically a ring of detectors is used in order to detect gamma rays coming from each angle. Thus, a PET scanner is typically substantially cylindrical to be able to capture as much radiation as possible, which should be, by definition, isotropic. Most modern PET scanners are composed of several thousand individual crystals (i.e., scintillator elements), which are arranged in two-dimensional scintillator arrays that are packaged in modules with photodetectors to measure the light pulses from respective scintillation events. The relative pulse energy measured by the photodetectors is used to identify the position of the scintillation event. Typically scintillator elements have a cross section of roughly 4 mm×4 mm. Smaller or larger dimensions and non-square sections are also possible. The length or depth of the crystal will determine how likely the gamma ray will be captured, and typically ranges from 10 to 30 mm. One example of a scintillation crystal is LYSO (or Lu1.8Y0.2SiO5:Ce or Lutetium Orthosilicate), which is chosen for its high light output, fast rise time, fast decay time, high average atomic number, and high density. Other crystals can be used.
Using Anger logic and crystal decoding, the source of each scintillation event can be identified as originating from a particular scintillator. A scintillation event will generate light initially radiating isotopically. The spatial distribution of this light may be modified by interactions with scintillator surfaces and reflectors before being detected by the four nearest photodetectors. From the relative pulse energy measured by each of these four photodetectors, the position of the scintillation event relative to the four photodetectors can be determined. The formulas for deriving position information from the relative pulse energies of the photodetectors are referred to as Anger arithmetic. These positions can be further refined by generating a lookup table from a floodmap in order to assign each scintillator event to a scintillator element using the lookup table. This process of mapping from the x- and y-positions obtained using Anger arithmetic to discrete scintillator elements is referred to as crystal decoding.
Often clinical PET data/images use only a relatively small number of gamma ray photons resulting from the use of fast-decay isotope and short scan times. Thus, the PET reconstructed images from these limited data sets appear noisy when the routine iterative reconstruction algorithms are applied. As a result, regularization, in-iteration and/or post filter smoothing methods are commonly used on the PET reconstruction images to improve image quality before a diagnosis. Improved and robust methods of reducing noise in PET images are desired.